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Electrospinning, a remarkable and versatile technique has been related to medical textiles, aiming to produce nanomaterials for drug delivery and tissue regeneration applications. Furthermore, electrospun nanofibrous materials with unique properties as favorable pore size distribution, porosity, surface area, and wettability, along with effective mechanical properties, are the frontrunner solutions. Also, the features of the nanofibrous structures can be designed and optimized by controlling electrospinning parameters related to the solution properties, the setup parameters, and the environmental conditions to design nanofibrous textile materials for the desired applications. Further, to accomplish the required functionality of the drug-delivery systems, a rather broad range of drugs have been loaded into the nanofibers using different electrospinning techniques, namely the blending, side-by-side, coaxial, tri-axial, emulsion, and multi-needle electrospinning, in order to accomplish specific drug-release profiles of the designed nanofibrous textiles. Thus, this chapter describes the different electrospinning techniques that have been utilized in the production of the textile nanofibrous materials as the application of these materials in bone, nerve, periodontal, and vascular regeneration, as well as in wound dressings, personal-protective-equipment (PPE), and cancer treatment, providing an overview of the recent studies and highlighting the current challenges and future perspectives for their medical applications.
FibEnTech Research Unit, Faculty of Engineering, University of Beira Interior, Covilhã, Portugal
Cláudia Mouro
FibEnTech Research Unit, Faculty of Engineering, University of Beira Interior, Covilhã, Portugal
*Address all correspondence to: igouveia@ubi
1. Introduction
During the last years, the huge and rapid development in nanotechnology has greatly contributed to the production of new functional materials with enhanced properties to be effectively and efficiently used in numerous fields [1, 2, 3]. Nevertheless, the development of medical textiles through the intervention of nanotechnology has been highlighted, in order to improve the health and quality of life of patients [3, 4].
Medical textiles, due to their distinctive features, such as their flexibility, wide range of sizes, and light-weight, physical, structural, and surface properties, have been produced with different materials (synthetic and/or natural) and with various structures (woven, knitted, braided, nonwoven, and composites) for different purposes, namely for advanced biomedical applications, like drug-delivery systems for tissue regeneration and wound healing [4, 5, 6]. However, developing advanced biocompatible, biodegradable, nontoxic, and nonallergic materials with desirable properties for drug loading and delivery in biomedical fields is still one the major challenges faced by the researchers, and thus, to minimize this gap, textile-based drug-delivery materials have been designed [3, 6, 7].
Among the different types of medical textiles, electrospun textiles fabricated from natural polymers, like proteins and polysaccharides, have gained increasing interest due to their low cost, biocompatibility, biodegradability, bioactivity, and unique properties, like the high surface area-to-volume ratio, smaller and more uniform fiber diameters, highly interconnected porous structures, and similarity to the extracellular matrices (ECMs) of tissues in the body [3, 8, 9, 10, 11]. Moreover, electrospinning has been revealed to be a simple and versatile method to produce drug-delivery systems, and hence the nanofibrous materials developed from the electrospinning are a unique and exceptional platform for loading and delivery of multiple therapeutic agents due to their performance and intrinsic nanoscale morphological characteristics [3, 8, 9, 10, 11]. Furthermore, the nanofibers drug loading has been achieved using different electrospinning techniques that comprise the incorporation of these therapeutic agents prior to the electrospinning process through the blended, coaxial, side-by-side, emulsion, and tri-axial electrospinning in order to deliver the appropriate amount of drug over the desired period of time and with a specific release profile [3, 5, 11]. Therefore, textile-based drug loading and delivery materials have been designed according to the medical conditions in order to improve the effectiveness of the drugs and reduce costs and the toxic side effects [5].
In the following sections of this chapter, an overview is provided which focuses on the fundamentals of the electrospinning technology, the different parameters that influence the diameter and arrangement of the produced nanofibers, and the main medical textile materials (e.g., cellulose, chitosan, collagen, alginate, keratin, and silk) used to produce textile-based drug loading and delivery nanofibers. In addition, the drug-release mechanisms and kinetics, as well as the different electrospinning techniques used to incorporate different therapeutic agents and produce electrospun nanofibers to act as drug-delivery textile materials, are described in detail. Finally, a brief outline of the recent studies concerning the production of textile-based drug loading and delivery materials for medical fields, namely for applications in bone, nerve, periodontal, and vascular grafts tissue engineering, wound dressings, and other textile-based materials, like personal protective equipment (PPE) and cancer treatment is presented, as well as the conclusions, challenges, and future perspectives of this emerging research field.
2. Electrospinning technology to produce textile-based drug-delivery materials
Electrospinning is a simple, versatile, and cost-effective manufacturing method that uses high voltage to produce an electrically charged jet of polymer toward the collector, where the continuous nanofibers are collected and interconnected nanofibrous structures produced (Figure 1) [3, 5, 8, 9, 12, 13, 14, 15]. Therefore, in the electrospinning process, a charged polymer jet is ejected when the electrostatic forces overcome the surface tension of the polymer solution and a Taylor cone-jet, subjected to a variety of forces, like an electric force imposed by the external electric field, a Coulomb force, a surface tension force, a viscoelastic force, and gravitational forces, is formed to produce the high-quality nanofibers [3, 8, 9, 12, 13, 14].
Figure 1.
Scheme of the basic device of electrospinning composed of four major parts: A spinneret, a syringe pump, a high-voltage source, and a collector.
In addition, the electrospun textiles have been recognized by their superior structural properties, like the high surface area-to-volume ratio, tunable porosities, morphologies, and diameters, as well as the ability to carry large amounts of drugs or to be chemically modified using wet chemical methods, oxygen plasma treatments, and in situ grafting polymerization for specific applications, such as drug loading and delivery materials [3, 5, 8, 9, 12, 14]. Moreover, there are various parameters that affect the morphology, diameter, distribution, and orientation of the produced textile nanofibers, namely the properties of the solution (e.g., viscosity, conductivity, and surface tension), the operating/process parameters (e.g., applied voltage, distance between the capillary tip and the collector, and the flow rate), and environmental conditions (e.g., humidity and temperature), and therefore, optimizing the variables that affect the electrospinning process is expected to develop textile nanofibrous materials with the specific features (porosity, wettability, mean diameter, and mechanical strength) desirable for tissue regeneration and wound-healing applications, as well as for PPE [3, 5, 8, 9, 12, 14].
Furthermore, the type of collector that is used in the electrospinning process determines the alignment of the electrospun nanofibers. Thus, using a stationary collector, the jet of the polymeric solution is directed to the collector, the solvent evaporates, and a dry and randomly oriented nanofibrous structure is obtained, mimicking ECM’s three-dimensional structures [3, 5, 12, 14]. On the other hand, using rotating collectors (e.g., discs, rotating drums, and rotating wire drums), nanofibers are deposited with a specific orientation and, consequently, are more prone to medical textiles for nervous and muscle tissue regeneration. Also, when the collector rotation speed is increased, the nanofibers diameter reduces, and the alignment improves [3, 5, 12, 14].
Additionally, electrospun textiles have also been produced from a broad range of polymer materials. Among them, synthetic polymers approved by the US Food and Drug Administration (FDA), such as polycaprolactone (PCL), poly(L-lactic acid) (PLLA), polylactic-co-glycolic acid (PLGA), polyvinyl alcohol (PVA), polyethylene oxide (PEO), and polyurethane (PU) have been explored for drug loading and delivery and biomedical applications due to their biocompatibility, excellent thermal stability, mechanical properties, and proper biodegradation rates [5, 8, 10, 12, 16]. Nevertheless, more recently, researchers started to explore natural biopolymers, such as proteins (e.g., hyaluronic acid, collagen, and silk fibroin) and polysaccharides (e.g., chitosan, alginate, and gelatin) according to their inherent nontoxicity, biocompatibility, bioactivity, sustainability, and eco-friendliness, as well as blending of synthetic and natural polymers to design electrospun nanofibrous textiles with notable medical properties [5, 8, 10, 12, 16].
Besides, textile nanofibers have been successfully electrospun and loaded with different drugs, including water-soluble, water-insoluble, anticancer, and antibacterial drugs through distinct techniques (see Section 3) in order to protect these agents from decomposition within the body before reaching the target site and produce textile-based drug loading and delivery nanofibers with varied drug-release profiles and kinetics, such as biphasic release, prolonged release, immediate release, and targeted release [3, 12]. These strategies can also control the incorporation of multiple agents in order to achieve the desired therapeutic effect and reduce the number of dosing times, improving patient compliance [3, 11, 12]. In this way, the electrospun nanofibers produced from biodegradable and biocompatible polymers using different electrospinning techniques have received increasing attention as carrier materials for drug loading and delivery in medical fields due to their flexibility, effectiveness, and unique physicochemical properties. Table 1 presents the different types of biocompatible synthetic and natural polymers that have been electrospun into nanofibers with controlled-release functions.
The most common biocompatible polymers used as drug-delivery carriers.
2.1 Mechanisms and kinetics for drug release
Textile nanofibers produced through the electrospinning process have been highlighted in the development of new nanomaterials, being able to provide improved drug-delivery systems with immediate- or prolonged-release profiles, minimum toxicity, and reduced dosage frequency due to their intrinsic properties (e.g., porosity, wettability, fiber diameters, and specific orientation), as described above [3, 5, 8, 10, 11, 17]. Besides, textile-based drug loading and delivery nanofibers fabricated from different types of polymer-carriers have been proposed to accurately predict the diverse drug-release kinetics from the polymeric matrix since the drug can interact with the polymer carriers in several manners and be released from nanofibers based on drug diffusion from pores, drug desorption from the surface layer, and/or polymer matrix degradation. Hence, different release mechanisms will be achieved depending on the characteristics of the drugs and the produced nanofibers (Table 2) [3, 5, 8, 11, 12, 17].
Drug-related factors
Effect on the drug-release kinetics
Drug loading amounts
Higher drug loading often results in faster release rates due to considerable amounts of surface-associated drug and the surface area-to-volume ratio of nanofibers.
Drugs in crystalline form are dispersed on nanofibers’ surfaces, and a burst release is observed, while the drugs in the amorphous form are deposited deeper inside the nanofibers and released in a sustained manner.
Solubility of the drugs in the polymers used as delivery carriers
Generally, the drugs highly soluble in the polymer matrix are released slower.
Drug-Polymer interactions
Physical and chemical interactions between the drugs and the polymer matrix result in a slower release, while the direct incorporation of the drugs in the nanofibers is associated with an undesirable burst release.
Nanofibers-related factors
Effect on the drug-release kinetics
Molecular weight of the polymer
High molecular weight polymers exhibit slower drug-release rates from nanofibers.
Nanofibers’ alignment
Randomly oriented fibers are often associated with quicker drug release due to the improved water uptake capacity.
Nanofibers’ diameters
Higher fiber diameters result in extended-release rates since the drugs diffuse in a greater space before reaching the edge of the nanofibers. In turn, nanofibers with smaller diameters get a quicker release due to their higher surface area available and dissolution rate.
Nanofibers’ crystallinity
Crystalline regions of the polymer are associated with a slower release when compared to amorphous domains.
Nanofibers’ porosity
The high porosity exhibited by the nanofibers might increase liquid absorption and, therefore, result in a faster drug release. Nonetheless, the nanofibers’ hydrophilicity, the pores sizes, and the volume of the pores influence the diffusion of the liquid which is absorbed by the nanofibers. Accordingly, it is frequently observed that thicker nanofibers with very high porosity release drugs faster in comparison to thinner fibers with low porosity.
Nanofibers’ surface area-to-volume ratio
Nanofibers with a high surface area-to-volume ratio display a greater space for interact with the nearby liquid and, consequently, are observed to have a quicker drug release.
Electrospinning techniques for nanofibers production
The drugs can be loaded in different layers of nanofibers through different electrospinning techniques. The choice of the technique allows to control the drug location in the nanofibers and the drug-release kinetics from the polymers.
Table 2.
Effect of the drug and nanofibers-related factors on the release kinetics.
For example, occasionally, it is desirable that the incorporated drug has the ability to produce rapid effects, being released in an immediate form. In these cases, a faster rate of drug diffusion from polymer structure should be achieved by using water-soluble and biodegradable polymers, like PVA, PEO, and Polyvinylpyrrolidone (PVP), being largely influenced by several factors, such as polymer swelling, wettability, porosity structure, polymer erosion, and drug dissolution/diffusion and distribution [3, 5, 8, 11, 12, 16]. On the other hand, a prolonged drug release, also denoted as a controlled, extended, and sustained release, requires the use of polymers or polymeric blends with a gradual degradation profile [3, 8, 11]. In addition, the electrospun nanofibers with an extended drug-release profile should display an adequate wettability and a thickness appropriate, as well as exhibit a core–shell structure containing a drug-loaded layer and an outer polymer layer that acts as a rate-controlling barrier. Moreover, core–shell nanofibers could provide biphasic drug-release profiles, which consist of an initial burst release followed by a sustained release (Figure 2) [3, 8, 11].
Figure 2.
Classification of the electrospun drug-loaded nanofibers based on their release profiles.
More recently, pulsatile drug-release profiles, where a sharp burst release is observed after a lag phase with no release, have gained increasing interest in various drugs or therapies. Besides, stimulus-responsive drug-release profiles have also been explored, and in these cases, responsive polymers after exposure to different stimuli, like pH, light, temperature, water, and CO2, can change their physicochemical properties (Figure 2) [3, 5, 8, 9, 11].
Therefore, the solubility of the drug in the polymeric solution and the structure of the drug-loaded fibrous matrix can give valuable information about the release kinetics and consequently allow to tune of the desired behavior [3, 5, 8, 9, 11]. Concerning that, controlling the multiple electrospinning parameters that affect the nanofiber’s features, such as fiber diameter, morphology, porosity, and wettability, and selecting the most suitable electrospinning technique for drug loading and subsequent controlled release can produce good-quality textile-based drug loading and delivery materials [3, 5, 8, 9, 11]. In addition, the choice of the drug molecules and methodology applied to incorporate the drugs into the textile electrospun nanofibers can favor the location, amount, and timing of drug release needed to modulate the release profiles and achieve the desired effects [3, 5, 8, 9, 11].
3. Electrospinning techniques for drug loading and delivery
A wide range of drug molecules, such as antibacterial molecules (e.g., antibiotics, nanoparticles (NPs), and natural extracts), molecules with biological activities, including anti-inflammatory, growth factors, vitamins, enzymes, and other proteins, as well as anticancer drugs, and nucleic acids (DNA and RNA) have been loaded into electrospun textile nanofibers through different electrospinning techniques, that comprise needle-based electrospinning and needleless electrospinning systems (Figure 3), in order to predict the location of drugs in the nanofibers and control the drug-release profile [3, 5, 8, 10, 11, 12, 16]. The drug loading using needle-based electrospinning techniques, like blending, side-by-side, coaxial, tri-axial, emulsion, and multi-needle electrospinning, add the drugs to the polymer solution before the electrospinning process and use a needle-like spinneret [3, 11, 12, 15, 18, 19]. On the other hand, needleless electrospinning methods, like the needle-free Nanospider technology, allow that multiple Taylor cones to be spun at the same time using a rotating electrode instead a needle, and hence this methodology is highly productive [3, 11, 12, 14, 18, 19]. In this way, the selection of the electrospinning technique can affect the role of the drug for its intended uses and manipulate its release rate and profile from the textile nanofibers [3, 11, 12, 14, 18, 19]. To accomplish that, the most common electrospinning methodologies for loading the drugs into electrospun textile nanofibers are described in detail below and their main advantages and disadvantages are summarized in Table 3.
Figure 3.
Schematic representation of several electrospinning techniques for the preparation of drug-loaded nanofibers.
Table 3.
Advantages and disadvantages of the main electrospinning techniques used to produce textile-based drug loading and delivery materials.
3.1 Needle-based electrospinning
The needle-based electrospinning techniques would have a direct impact on the drug loading and release behavior from the electrospun nanofibers. The needle-based systems can result in simple, core-shell, and multi-axial nanofibers comprising more than one textile material [3, 11, 14, 18]. However, in the needle electrospinning techniques, capillary tubes, such as a needle-like spinneret, are required to form the nanofibers [3, 11, 14, 18].
3.1.1 Blending electrospinning
Blending electrospinning is the simplest and most basic method that uses a single nozzle to incorporate drugs into the electrospun nanofibers by blending them into the polymeric solution before the electrospinning process [14, 15]. In addition, different types of carrier materials and drugs can be processed by blending electrospinning, and hence it is the most researched and applied strategy for loading drugs into electrospun nanofibers. However, when this technique is used, in general, the drugs loaded are rapidly released from the nanofibers due to their homogeneous distribution all over the surface of the polymeric nanofibers and their high surface-to-volume ratio that favors the drug release. Nonetheless, their release depends on the degree of drug encapsulation into the polymeric matrix and the affinity between the drug and the polymer [14, 19]. Thus, to avoid an initial burst effect and ensure a more sustained release, the drugs absorbed and/or encapsulated into nanostructures, like NPs, nanospheres, nanomicelles, and nanotubes, can be added to the polymeric solution before the electrospinning process [14, 19].
Moreover, drugs with different properties can be blended into the same polymer carrier, and nanofibers with a biphasic drug release are obtained. For example, Li et al. dissolved PLGA (25% (w/v)) in a mixture of acetone and N,N-dimethyl formamide (DMF) with a volume ration of 3:1, and dexamethasone (DEX) at a concentration of 15% and green tea polyphenols (GTP) at concentrations of 5%, 10%, and 15% by weight of PLGA were added to the blend [22]. Afterward, the polymeric solutions were electrospun at a flow rate of 0.5 mL/h, using a working distance of 18 cm and an applied voltage of 23 kV. The results revealed that the DEX and GTP released from the electrospun nanofibers in PBS at pH 7.4 achieved a biphasic release profile. The hydrophilic GTP drug exhibited a faster release, while the hydrophobic DEX was successfully released from the channels formed by the fast-released GTP molecules. Thus, drugs with different hydrophilicity were used to achieve a biphasic release profile from the same polymer solution, which was controlled by the drugs’ diffusion rate and the degradation of the polymer carrier [22]. Furthermore, stimuli-responsive nanofibers could also be produced by blending electrospinning. Li et al. produced electrospun silk fibroin/PEO (SF/PEO) nanofibers with a pH-responsive controlled-release behavior [23]. For this purpose, 5 mol% strontium or copper-doped hollow bioactive glass nanospheres (5Sr/Cu-HBGNs), synthesized via a sol–gel method, were loaded with a 2.5 mg/mL aqueous solution of vancomycin hydrochloride (VAN), a glycopeptide antibiotic that inhibits bacterial cell wall synthesis, and incorporated in a mixture of 10% SF/3% PEO to produce 2% w/v solutions. The release kinetics behavior of VAN from VAN@5Sr/Cu-HBGNs/PEO/SF nanofibers in PBS at pH of 4.5 and 7.4 demonstrated that VAN@doped HBGNs/PEO/SF nanofibers presented a slower and sustained release rate in comparison with VAN@HBGNs/PEO/SF. Besides, the cumulative release of VAN from the electrospun VAN@5Sr/Cu-HBGNs/PEO/SF nanofibers at pH 7.4 was faster than at acidic microenvironments (pH 4.5), indicating that the drug release increased with an increase in pH along with a quick VAN desorption release. Thus, their pH sensitivity can be explained by the different solubility and both hydrophilic nature and amphoteric properties of SF under acid and alkali conditions [23].
3.1.2 Side-by-side electrospinning
The side-by-side electrospinning technique is a variation of the basic single nozzle electrospinning comprised of two capillaries placed one adjacent to the other to produce nanofibers with Janus structures and materials with two distinct layers [3, 14, 18]. However, this method is not easy to perform, since both capillaries, containing different polymer solutions, are controlled by the same pump, and the same voltage is applied to both solutions, which causes repulsion between the two different polymers and separate from each other [3, 14, 18]. Hence, few studies have been developed in order to fabricate electrospun textile nanofibers with a Janus structure [3]. However, the side-by-side bi-component fibers exhibit the characteristics of both polymers, and consequently, we can achieve a biphasic drug-release profile required for a particular type of drug and application [14]. Thus, a biphasic drug release can be achieved by incorporating the drug on both sides of the fiber, where one side consists of a water-soluble polymer and the other side of a water-insoluble polymer [14]. On the side composed of the water-soluble polymer, the drug diffusion rate will be affected by polymer erosion and will be faster, while on the side containing the insoluble polymer, the drug release will be extended [14].
Shi et al. used side-by-side electrospinning to combine a solution of 10% (w/v) PLGA with an anti-inflammatory drug (0.2% (w/v) valsartan, V) and a solution of 15% (w/v) PVA with copper sulfide nanoparticles (CuS NPs) and an antibacterial drug (mupirocin, M) [24]. Both solutions were placed in two different syringes with 26 G blunted spinning needles and electrospun at a constant feeding rate of 1.0 mL/h, using a working distance of 20 cm and an applied voltage of 15 kV. The Janus amphiphilic nanofibers composed by a hydrophobic side of PLGAV-CuS NPs and a hydrophilic side of poly(vinyl alcohol) mupirocin nanofiber (PVAM) exhibited a biphasic drug release mechanism. The cumulative release of the hydrophilic antimicrobial drug, mupirocin, from PLGAV-CuS/PVAM nanofibers enabled a continuous and slow release, reaching 89.36% after 24 h of incubation [24]. The sustained release of mupirocin can be attributed to the hydrogen bonding between the drug and the PLGAV-CuS/PVAM. In addition, the water uptake ability of the PLGAV-CuS/PVAM nanofibers could have a certain blocking effect on the release of mupirocin. In turn, the release profile of valsartan reached only 23.77% after 24 h, providing a slow release of drug from nanofibers. This result confirms that the diffusion of hydrophobic drugs in a hydrophilic environment is hindered. Therefore, the PLGAV-CuS/PVAM Janus nanofibers initially released mupirocin more quickly, and then, the valsartan and remaining mupirocin were continuously released, achieving tunable antibacterial and anti-inflammatory gradient drug release systems [24].
3.1.3 Coaxial electrospinning
Coaxial electrospinning technique also uses two capillaries, like side-by-side electrospinning, however, in coaxial electrospinning the capillaries are arranged one inside the other. In addition, the concentric capillaries, i.e. the inner capillary coaxially placed inside the outer one, are connected to two independent reservoirs, with controllable flow rates [11, 14, 18]. Concerning that, coaxial electrospinning allows the production of core-shell nanofibers with different core and shell compositions, including functional fibers encapsulated with different drugs and hollow nanofibers with adjustable shell thickness by selectively removing the core from as-spun core-shell nanofibers [11, 14, 18].
Hence, core-shell nanofibers produced from coaxial electrospinning provide an added advantage as carriers for drug delivery, protecting the drugs’ native structure and their bioactivity from harsh environments during textile nanofibers production [11, 14, 19]. In addition, the drugs loading into the core of the textile nanofibers can significantly reduce and/or prevent the initial burst release rate from the polymeric matrix and maintain a sustainable drug release for an extended and controlled time [11, 14, 19]. Moreover, this technique contributes greatly to the high loading capacity of diverse drugs and allows susceptible drugs to be delivered. Furthermore, core-shell nanofibers prepared from miscible and immiscible polymers can be produced through coaxial electrospinning, as well as from nonspinnable solutions and polymers and drugs with low compatibility [11, 14, 19]. However, this technique requires a special apparatus, and the operating conditions should be carefully chosen to ensure desirable results [11, 14]. Nevertheless, coaxial electrospinning has the ability to adjust the thickness of both layers by changing the flow rate, particularly the shell thickness that acts as a barrier to diffusion, controls the release kinetics of the encapsulated drugs, and enables the potential to simultaneously load various components [14, 19]. Therefore, core-shell nanofibers fabricated from the coaxial electrospinning technique can effectively avoid the initial drugs’ explosive release and achieve a sustained release profile. For example, Li et al. successfully produced Xylan/PCL core–shell nanofibers using coaxial electrospinning. In this study, Xylan aqueous solution (20 w/w%) containing the anti-inflammatory drug levofloxacin (LEV) was used as core and coated by (12 w/w%) PCL solution (shell). The solutions were placed in two coaxial injectors and electrospun at an applied voltage of 14 kV, using a collecting distance of 15 cm, a constant flow rate of 1.0 mL/h for the Xylan solution and a flow ration between Xylan and PCL of 1:1.1, 1:1.2, and 1.1.4, respectively. The results obtained revealed that the hydrophobicity of the PCL prevents a burst release and short diffusion of the LEV loaded into the water-soluble Xylan core. Additionally, the increase in PCL’s flow rate resulted in a lower degradation rate of the nanofibers due to the thicker fibrous shell layer and consequently can be used to produce a slow drug release [25].
3.1.3.1 Tri-axial electrospinning
Coaxial electrospinning can be used to produce textile nanofibers with more complex architectures, namely by using a tri-axial electrospinning technique, composed of three concentric capillaries [3, 14]. The three-layer core-shell nanofibers are formed from an external polymeric solution, an intermediate polymeric solution, and a core polymeric solution controlled by different pumps [3, 14]. Besides, the tri-axial nanofibers can be produced with the core enclosed by two different polymer layers or with two layers of the same polymer. Moreover, the core polymers can also be surrounded by an outer polymer layer and a layer of void areas. In turn, different drugs can be incorporated into the core, or in each different layer, and released by diffusion and degradation [3, 14]. The three-layer textile nanofibers can be also prepared with targeted drug-release functions for medical applications in order to improve the drug efficacy and reduce its toxicity and side effects, as well as with a linear and constant release rate in order to eliminate the drug burst release [3, 14]. Likewise, the drug content can be changed by using a tri-axial electrospinning technique, creating an increased drug gradient distribution from the outer layer to the inner layer. Thus, in comparison with the double-layer core-shell nanofibers produced from coaxial electrospinning, the tri-layer core-shell nanofibers prepared by tri-axial electrospinning provide more potential to overcome the limitation of poor drug solubility, protect drugs from the adverse environments, and control the drug-release kinetics for developing functional textile-based drug loading and delivery materials [3, 14, 26].
Ding et al. compared the capacity of aspirin-loaded ES100 core-shell nanofibers (produced with an improved tri-axial electrospinning system) (CSFs) with the monolithic composite (prepared with the same material using a traditional blended single-fluid electrospinning) (MCFs) for a sustained drug release. The CSFs revealed better results than the MCFs, although the release of aspirin from both occurred through an erosion mechanism [27]. The CSFs exhibited a prolonged release of the drug and showed a lower release rate during the first 2 h under acidic conditions. In addition, in a posterior neutral environment, a prolonged drug release was obtained by an extended-release effect [27].
3.1.4 Emulsion electrospinning
Emulsion electrospinning is a uniaxial electrospinning technique, which has attracted growing interest in the production of core-shell textile nanofibers without the use of specific coaxial apparatus [11, 14, 18, 19]. This method is similar to conventional electrospinning, apart from that the polymeric solution is replaced by a water-in-oil (W/O) or oil-in-water (O/W) emulsion [14]. In the emulsion electrospinning, most emulsions are of the W/O type, where the drug is usually dissolved in an aqueous phase composed mainly of water-soluble polymers (i.e., the water phase) and then dispersed in a continuous-phase composed of an organic polymer solution (i.e., the oil phase) [14, 19]. The droplets of the water phase dispersed in the oil phase evaporate more slowly than the organic polymer solution, resulting in a viscosity gradient between the two different phases. Subsequently, this gradient guides the droplets of the water phase from the surface to the center, the droplets are stretched into elliptical shapes along the axial region under a high voltage and the core-shell nanofibers containing the drug into the core are formed [14]. Hence, W/O emulsions, in which the hydrophilic drugs are dissolved in an aqueous polymer solution, while hydrophobic polymers are dissolved in organic solvents, are particularly useful for a drug continuous and sustained release, avoiding the initial burst release, typical of most drug-delivery materials, as well as play a role in protecting the drugs in the core from the harmful effects of the external environment, enhancing their bioactivity and effectiveness [11, 14, 19]. Moreover, the drugs are released faster in the shell than in the core layer; once the drugs are loaded into the core have to pass through the core-shell matrix before being released. Furthermore, good-quality core–shell nanofibers can be produced through emulsion electrospinning using diluted polymer solutions [19]. Various combinations of hydrophilic drugs and hydrophobic polymers with low compatibility and affinity can also be explored, removing the necessity of using a common solvent for both the drug and the polymer [18]. However, emulsifiers such as surfactants are frequently used to stabilize the emulsions and the encapsulated drugs [11, 19].
Weng et al. prepared camellia oil-loaded zein nanofibers with a core-shell structure using an emulsion electrospinning technique. For this purpose, zein (30% v/v) aqueous solution was firstly prepared in acetic acid (70% v/v) and then camelia oil was dropped at 10, 20, 40, and 60% (v/v) based on the zein solution. After stirring for 1 h, 30% (w/w) of glycerol (based on the weight of zein) was added as a plasticizer. The blends were mixed with a high-speed homogenizer at 1000 rpm for 3 min and then treated by ultrasonic for 2 min at 250 W at 25°C to obtain the required O/W emulsions. Afterward, the emulsions were placed in the syringe and electrospun at a flow rate of 0.5 mL/h, using a working distance of 12 cm and an applied voltage of 20 kV. The results of this study demonstrated that core–shell nanofibers of zein with camellia oil could be used as a promising controlled-release carrier for hydrophobic bioactive compounds [28].
3.1.5 Multi-needle electrospinning
The multi-needle electrospinning is one of the simplest techniques to achieve nanofibers with high throughput [20]. This electrospinning technique implies passing a polymeric solution through multiple needles connected to a high-voltage supply. In addition, the flow rate is controlled by a single pump and is required the application of a higher voltage to continuous electrospinning due to the large mass of the spinning solution [20]. Nevertheless, multi-needle electrospinning presents several drawbacks, such as unstable electric field strength, changes in fiber size distribution, clogging at the tip of the needles, and cleaning of multiple needles. In addition, the multi-needles may also provide repulsion from adjacent jets. Nonetheless, multiple drugs can be added, and diverse release profiles can be achieved [20].
Varesano et al. tested several multi-jet electrospinning setups using a 7 wt% PEO aqueous solution and varied the number of nozzles between 2 and 16. The data obtained showed that when increasing the number of jets above 6 was imperative to increase the collector dimensions because the deposition area also increased [29]. Hence, multi-needle electrospinning systems have great potential for large-scale nanofibers production. Moreover, Yoon et al. fabricated Polystyrene (PS)/Polyamide 6 (PA6) nanofibers using multi-jet electrospinning and revealed that the fiber content was directly related to the syringes’ number used for PS and PA6, respectively [30].
3.2 Needleless electrospinning techniques
The needleless electrospinning does not use a needle-like spinneret but static or rotating spinnerets, which can be a cylinder, a ball, a disc, and/or a rotating electrode, immersed into an open container with the polymeric solution in order to produce nanofibers in higher quantities for industrial range [18, 20]. Nevertheless, when the open container is filled with a polymer solution dissolved in volatile solvents, the solvent evaporates quickly and can affect the reproducibility of fiber morphology. Moreover, in needleless electrospinning, many Taylor cones (the source of nanofibers) are created simultaneously on the surface of the spinnerets, and hence these techniques are highly productive and more effective in producing high-quality nanofibers [20, 21]. For example, in the Nanospider technology, a needleless electrospinning technique, a rotating electrode/roller transport the polymer solution from the container, allowing industrial-scale production of nanofibers [20, 21]. In this way, Nanospider is a pilot electrospinning equipment, already used on an industrial scale, which allows an easy scale-up. Furthermore, this is a simple and versatile technique for textile nanofibers production from a wide range of materials [20, 21]. Regarding that, Mouro et al. incorporated Hypericum perforatum L., a medicinal plant extract, in a PLLA/PVA/CS emulsion via Nanospider, a needleless electrospinning technology. The results obtained showed a controlled and sustained release profile of the Hypericum perforatum L. from the produced nanofibers for 72 h. In addition, the release speed was dependent of the content of plant extract incorporated and the swelling of the produced nanofibers [31].
Recently, electrospinning derivative technologies, such as melt electrospinning and centrifugal electrospinning, have also arisen as promising approaches [3, 8, 14, 32, 33]. The melt electrospinning is a solvent-free process and hence a more efficient and environmentally friendly technique. A typical melt electrospinning setup includes a heated polymer fed into a spinneret. However, like traditional electrospinning, a high electric potential is applied between the spinneret and the collector for producing the textile nanofibers [8, 14, 33]. Moreover, melt electrospinning enables the production of nanofibers in the absence of residual solvents because it does not require the dissolution of the polymers in organic solvents. In addition, various materials can be used to fabricate nanofibers. However, this technique only uses materials that melt, and polymer degradation or denaturation can occur because high temperatures are applied [8, 14, 33]. On the other hand, the centrifugal electrospinning technique combines the electrostatic force of the electrospinning and the centrifugal force to promote jet stretching [3, 32]. Centrifugal electrospinning is a simple and cost-effective method that allow the fabrication of highly aligned and small-diameter textile nanofibers, upon applying adjusted rotational speeds and voltages, for drug delivery and regenerative medicine applications [3, 32]. Regarding that, in centrifugal electrospinning, the electrical fields allow the spinneret to rotate at a slower speed and simultaneously a lower voltage to overcome the surface tension of the spun solution is required [3, 32].
4. Applications of the electrospun textile drug-delivery materials in medical fields
The unique features of the electrospun textile materials, namely the high surface area, tunable porosity, and smaller diameter of the fibers, as well as the easy optimization of the nanofiber’s properties through controlling the different parameters, like the solution properties, processing parameters, and environmental conditions, made them ideal for drug loading and delivery applications [3, 5, 8, 9, 12, 14]. Moreover, the various electrospinning techniques allow for incorporating a wide range of drugs with high loading and encapsulation efficiencies into diverse polymer-textile nanofibers [3, 5, 11, 12, 14, 18, 19].
Therefore, this section is focused on the textile-based drug loading and delivery of nanofibrous materials produced from the different electrospinning techniques for medical applications. Regarding that, it is provided an overview of the recent studies for applications in bone, nerve, periodontal, and vascular regeneration, as well as in wound dressings and other textile-based materials, like PPE and cancer treatment. Furthermore, Table 4 are presented several studies where the authors prepared drug-loaded nanofibers using different electrospinning technologies for medical applications.
Medical application
Drugs/biological agents
Carrier/polymer
Drug-release highlights
Reference
Wound treatment
Amoxicillin (AMX)
Pullulan (Pull) and poly(lactide-co-glycolide) (PLGA)
The release rate of the drug from AMX-loaded Pull nanofibers was fast. In turn, when the AMX-loaded Pull nanofibers were sandwiched between extra electrospun PLGA layers was observed a continuous sustained release favorable for wound-healing dressings. The PLGA layers around the AMX-loaded Pull work as a diffusion barrier, decreasing the release rate of AMX.
The sandwich-like PCL/Gelatin/PCL electrospun multilayered mats were incorporated with CTZ. The release profile of CTZ-loaded Gelatin from the internal layer was found to be dependent on the PCL structure, which stabilized and trap the CTZ, avoiding a fast initial burst. The result was a continuous and slower sustained release profile.
Polyamide-6 (PA6) and Hydroxyethyl cellulose (HEC)
Three different ratios of TA (5, 10, and 25% wt, versus the total weight of polymer) were incorporated into the polymer blend solution. In vitro release of TA from PA6/HEC nanofibers displayed no burst release. The PA6/HEC/25%TA nanofibers exhibited a stable TA release rate that reached 80% after 24 h of incubation. Thus, the release rate has a direct relation with the TA concentration (up to 25%).
The electrospun nanofibers loaded with TMB, a hemostatic agent, and VCM, an antibacterial agent, were placed in PBS pH 7.4 for 72 h and revealed a burst release within the first 2 h followed by a sustained release, indicating its potential as a dressing material for hemostatic wound-healing applications.
The Nrg was loaded into PCL/PEG nanofibers with various amounts (3.125, 6.25, 12.5% Wt. Nrg to polymer) and the release profiles revealed a burst at the first 2 h. Then, a steady release was observed, reaching approximately 70, 42, and 19% within 24 h for PCL/PEG/Nrg (3.125%), PCL/PEG/Nrg (6.25%), and PCL/PEG/Nrg (12.5%) NFs, respectively. So, interestingly the release rate was faster in the nanofibers with low Nrg ratio.
1 and 10% of Nimodipine, a calcium-antagonistic drug with preferential cerebrovascular action and neuroprotective properties, was successfully loaded into PLGA nanofibers. The results indicated a controlled and sustained drug release over 4–8 days, highly dependent on release conditions. Therefore, the electrospun Nimodipine-loaded PLGA nanofibers are a promising approach for intracranial use.
The SF containing the CEM drug was embedded into PVA with Ag doped MSi NPs to produce three-dimensional sandwiched nanofibers. The prepared sandwich layered nanofibers exhibited a synergistic antibacterial effect due to the controlled and sustained release of CEM from the middle layer. No burst release was observed.
Vancomycin hydrochloride (Van), Simvastatin (Sim), and Fluorescein isothiocyanate-bovine serum albumin (FITC-BSA)
PCL and Gelatin
PCL/Gelatin co-electrospinning membranes were loaded with different model drugs (e.g. Van and FITC-BSA as hydrophilic drugs/proteins and Sim as a hydrophobic drug) in order to evaluate their potential for bone regeneration. The results showed that the presence of the PCL can avoid a fast degradation and dissolution of the Gelatin and consequently an early explosive release of the hydrophilic drugs/proteins. Thus, by increasing the PCL content extended drug release can be obtained.
PCL, Polyethylene oxide (PEO), and Hyaluronic acid (HA)
Core-shell nanofiber membranes were produced via coaxial electrospinning to promote tendon recovery posttendon injury. The nanofibers comprised a PCL shell and a core of PEO and different ratios of HA/Platelet-rich plasma (PRP) where were loaded different growth factors. The results showed that the release of growth factors from the core region of the core-shell nanofibers was linear for 10 days and followed a sustained release behavior after 35 days.
PVA, Trigonella foenum-graecum(TFG), and Coriandrum sativum(CS)
Electrospun composite nanofibers composed of PVA, TFG, and CS were incorporated with 5-FU, a water-soluble, hydrophilic antimetabolite drug. The composite nanofibers showed a sustained drug release over 4 h, emphasizing their suitability for ocular drug-delivery systems.
The paclitaxel-loaded liposome-incorporated Chitosan/PCL-Chitosan core-shell nanofibers (produced with a coaxial electrospinning system) do no demonstrate an initial burst release under Ph values of 5.5 and 7.4. Additionally, a linear release of paclitaxel from the liposome-loaded core–shell nanofibers was observed, making them effective for the treatment of various tumors.
Core–shell Gelatin/PLA nanofibers containing an anticancer DTX drug were obtained through emulsion electrospinning. The release profiles revealed that the cumulative DTX release was higher in acidic pH (6.5) than in neutral pH (7.4), enhancing the efficacy of drug delivery to tumor tissues in an acidic environment.
PCL/Gelatin nanofibers containing DOX-loaded amine-functionalized mesoporous silica NPs (MSNs), linked with green synthesized Ag NPs, revealed a proper release profile of DOX and Ag NPs for healing wounds caused by melanoma cancer.
A trilayered material was prepared with a PCL outer layer and encapsulated with microRNA-126 and microRNA-145 in the Arg-Glu-Asp-Val (REDV) peptide-modified PELCL inner and PELCL middle layers, respectively. The data obtained showed a fast release of microRNA-126 and/or a slow release of microRNA-145 suitable for replacement of artificial small-caliber blood vessels.
Several drug-loaded nanofibers using different electrospinning techniques for medical applications.
4.1 Bone tissue regeneration
Nanofibers produced from electrospinning have been investigated for bone regeneration due to their intrinsic properties, such as biocompatibility, highly porous structure, small pore size, high surface area, three-dimensional (3D) architecture, and mechanical features that are compatible with the natural nanostructure of bone. In addition, different drugs have been incorporated into these nanofibrous structures to improve bone regeneration [8, 16].
In 2022, Canales et al. produced electrospun fibers of Poly(lactic acid) (PLA) containing bioactive glass (n-BG) and magnesium oxide (n-MgO) NPs to be used in bone tissue engineering applications [49]. Their results showed that the electrospun PLA/n-BG and PLA/n-BG/n-MgO fibers presented a significant increase in fiber diameter with mean diameter values of 3.1 ± 0.8 μm, while the neat PLA and the PLA/n-MgO displayed an average fiber diameter of 1.7 ± 0.6 μm [49]. Besides, the electrospun PLA/n-BG/n-MgO fibers presented a high porosity and an interconnected pore structure required for tissue engineering applications. However, the incorporation of the NPs affected the thermal properties and mechanical properties [49]. All the composite fibers containing n-BG showed the capability to precipitate hydroxyapatite on the surface, demonstrating their bioactivity, while the addition of the n-MgO to PLA provides antibacterial properties against S. aureus [49]. Moreover, the osteoblastic phenotype expression ability of the produced composite fibers was assessed in comparison with the neat PLA fibers measuring the alkaline phosphatase expression (ALP), a marker of osteoblastic activity known for its capacity to promote cell differentiation, and the results indicated that the PLA/n-BG presenting the highest osteoblastic expression. Therefore, this study revealed that the incorporation of both n-BG and n-MgO NPs into the PLA can be considered to produce a synergic effect increasing its bioactivity and antimicrobial behavior [49].
Huang et al. encapsulated citrate-stabilized gold-nanoparticles (GNPs) into PVP/Ethylcellulose (EC) by coaxial electrospinning technique in order to create core-shell nanofibrous membranes [50]. In this study, the core-shell nanofibers were prepared by coaxial electrospinning with EC as the core and PVP containing GNPs as the shell. The GNPs-loaded electrospun PVP/EC nanofibers were prepared by changing the feeding GNPs to 0.5:1, 1:1, and 1.5:1 (P/E-0.5, P/E-1, and P/E-1.5), respectively [50]. The data obtained revealed that the GNPs were successfully encapsulated into the electrospun nanofibers. Besides, the GNPs addition did not significantly affect the morphology of the nanofibers, although it improved the porosity and the mechanical properties of the nanofibers [50]. Furthermore, the presence of the GNPs in electrospun PVP/EC membranes triggered a higher alkaline phosphatase activity, mineralized nodule formation, and osteogenic-related genes expression, thus confirming the excellent biocompatibility and osteogenic bioactivities of the produced composite material [50]. The in vivo studies also confirmed that GNP-incorporated electrospun PVP/EC nanofibers accelerate bone regeneration. Hence, the produced coaxial electrospun membranes demonstrated the appropriateness for being considered as a promising material for bone repair [50]. In a similar way, Alam et al. produced coaxial PCL/Gelatin/Poloxamer 188 (P-188) nanofibers for duel release of β-lactoglobulin (a hydrophilic protein) and vitamin K2 (a hydrophobic agent) from core and shell fibers [51]. The coaxial electrospun nanofibers produced from the PCL/P-188 shell and Gelatin/P-188 core displayed superior stability, average weight loss, no water uptake ability, and a sustained, controlled release of both β-lactoglobulin and vitamin K2 [51]. Moreover, these nanofibers also significantly enhanced alkaline phosphatase activity and promoted higher cell viability for human osteogenic sarcoma cells (Saos-2 cells), and thereby, coaxial PCL/Gelatin/P-188 nanofibers can release both hydrophilic and hydrophobic bioactive compounds and be a promising bioactive material for bone tissue engineering with improved osteogenic properties [51].
Electrospun nanofibers with a core-shell structure have also been produced by emulsion electrospinning for bone tissue regeneration. For example, Boraei et al. fabricated core-shell nanofibers from PVA and PCL by water-in-oil electrospinning encapsulated with strontium ranelate (SrR), an osteogenic agent [52]. The authors successfully obtained nanofibers with a shell of PCL (the oil phase) and a core of PVA incorporated with SrR (the water phase). In addition, higher contents of SrR induced the formation of fibers with increased diameters and decreased the crystallinity of the nanofibers [52]. Moreover, the SrR release from the electrospun core-shell nanofibers occurred through a Fickian diffusion mechanism, and it was more evident a quicker drug dissolution in the samples with higher SrR content [52]. Additionally, the incorporation of the SrR improved mesenchymal stem cell proliferation and enhanced the expression of ALP, Runx2, Col I, and OCN genes. The in vivo assays also showed that animal models treated with the core-shell nanofibers displayed an increased bone formation of calvarial defects [52]. Further, Al-Baadani et al. investigated the suitability of the electrospun PCL/Gelatin membranes containing different drugs/proteins, like fluorescein isothiocyanate-bovine serum albumin (FITC-BSA), vancomycin hydrochloride (Van), and simvastatin (Sim) to improve the antibacterial and osteogenic activities [42]. This approach combined the biocompatibility of the Gelatin, namely its ability to improve the adhesion and differentiation of osteoblasts, with the mechanical properties of the PCL. The results showed that the produced membranes from coaxial electrospinning could be utilized as drug-carriers for slow-release hydrophobic drugs, like Sim, loading in the PCL solution, and for controlled-release of hydrophilic drugs/proteins, like FITC-BSA and Van, loading in Gelatin solution, adjusting the PCL content [42].
Additionally, the electrospun nanofibers have also been applied to restore other structural tissues, namely skeletal muscle and cartilage ligament.
4.2 Nerve tissue regeneration
Neural tissue repair is one of the major challenges faced by the researchers because the most common neural injuries result in an irreversible loss of function [16]. In this context, biocompatible polymer nanofibrous conduits, with customized lengths and sizes and controlled delivery of drugs for nerve regeneration, have recently been investigated using different electrospinning techniques. In addition, both aligned and randomly oriented nanofibers have been further explored in order to improve neural stem cell adhesion and differentiation, as well as neural tissue regeneration [8, 16].
Alipour et al. explored the encapsulation of fingolimod-loaded PLGA nanoparticles into electrospun PU/PCL/Gelatin nanofibers for promoting neurite formation and axonal regeneration. Different amounts of fingolimod (0.01, 0.02, and 0.03%), a useful drug in nerve regeneration, were studied. The increased content of drug loaded into the membranes resulted in an increase in the nanofiber’s mean diameter [53]. Nevertheless, the nanofiber’s diameters of the produced membranes are in the acceptable range displayed by nerve tissues, as well as the degradation rate, the mechanical properties, and the water uptake behavior [53]. Besides, the electrospun PU/PCL/Gelatin nanofibrous membrane loaded with 0.01% fingolimod through PLGA NPs provided a more favorable environment for cell growth and proliferation. Moreover, the release profile of the fingolimod from this membrane displayed a burst release behavior during the first 24 h, and then a slow release occurred for up to 5 days [53]. Therefore, the electrospun PU/PCL/Gelatin nanofibers loaded with fingolimod NPs can be a suitable candidate for application in neural tissue engineering [53].
Xia et al. incorporated the recombinant human vascular endothelial growth factor (VEGF) and the recombinant human nerve growth factor (NGF) on the shell and in the core of the PLLA nanofibers by emulsion electrospinning [54]. The electrospun core-shell nanofibrous membranes with VEGF on the shell and NGF in the core displayed a sequential release profile, in which most of the VEGF was released in the first few days, and the NGF could be gradually released from the PLLA core nanofibers for >1 month [54]. Besides, the nanofibers containing the NGF and VEGF encouraged the nerve differentiation of induced pluripotent stem cells-derived neural crest stem cells (iPSCs-NCSCs) in vitro [54]. Moreover, in vivo data showed that, after 3 months, animals treated with the dual-delivery of VEGF and NGF membrane showed a significant improvement in neovascularization and nerve healing postoperation [54].
In turn, Alipour et al. reported the incorporation of vitamin C (VC) at concentrations of 5, 10, and 15 wt.% into PCL/Polyglycerol Sebacate (PGS) nanofibers [55]. The nanofibers produced from a rotating drum collector presented an aligned and uniform appearance, although the incorporation of the VC induced a decrease in the nanofibers’ mean diameter. In addition, the electrospun membranes exhibited appropriate properties for nerve applications, namely mechanical properties, wettability, water uptake capability, and degradation [55]. Furthermore, the release profile of VC was characterized by an initial burst, followed by a gradual release, which is fundamental for more beneficial therapeutic effects on peripheral nerve regeneration. However, the electrospun PCL/PGS nanofibers containing 5 wt.% VC provided a more favorable environment for PC12 cell adhesion and migration [55].
Chen et al. prepared multilayered membranes with a PCL outer layer, a Fe3O4 magnetic NPs (Fe3O4-MNPs)/PCL middle layer, and a melatonin (MLT)/PCL inner layer for peripheral nerve regeneration [40]. The results obtained revealed a sequential and sustainable drug release appropriate for mitigating oxidative stress and inflammatory response and inducing a microenvironment suitable for nerve regeneration [40]. In fact, the composite nanofibrous membranes showed satisfactory mechanical strength, in vitro biocompatibility in contact with rat Schwann cells (RSC 96), and beneficial effects for nerve regeneration in vivo [40]. Further, Mohamadi et al. obtained electrospun conduits from the PCL/collagen/nanobioglasses (NBG) blends, and the NGF was loaded in the structure of conduits [56]. Their results showed that the PCL/Collagen/NBG containing NGF revealed the potential to be applied to promote sciatic nerve regeneration. In addition, the NGF-loaded conduits exhibited the ability to enhance the recovery of the injured sciatic nerve [56].
In 2023, for the first time, Puhl et al. delivered mRNA encoding NT-3 from the aligned electrospun PLLA fibers for peripheral nerve regeneration [57]. In this approach, the PLLA fiber’s surface was functionalized with poly(3,4-dihydroxy-L-phenylalanine) (pDOPA) or coated with dextran sulfate sodium salt (DSS). Then, lipoplexes were obtained by complexing the synthesized pseudouridine-5’-triphosphate (Ψ)-modified mRNA encoding NT-3 (ΨNT-3mRNA) with the cationic delivery vehicle JetMESSENGER® and immobilized to the fiber PLLA surface [57]. The cationic lipoplexes containing ΨNT-3mRNA complexed to JetMESSENGER® were immobilized into PLLA fibers resulting in detectable ΨNT-3mRNA release for 21 days. Moreover, the ΨNT-3mRNA/JetMESSENGER® lipoplex-immobilized pDOPA-coated aligned electrospun PLLA fibers showed the capability to support Schwann cell secretion of NT-3 and enhance neurite outgrowth from dorsal root ganglia (DRG) neurons [57].
4.3 Periodontal tissue regeneration
Periodontal tissue is a complex tooth-supporting connective structure, in which soft, mineralized connective and epithelial tissues are structured to form a dentogingival junction [8]. However, it is highly susceptible to chronic periodontal diseases, like periodontitis, resulting in irreversible tissue destruction. Thus, to restore tissue integrity and improve periodontal regeneration, researchers have produced diverse nanofibrous membranes via electrospinning and incorporated different drugs with antibacterial, anti-inflammatory, and tissue regeneration properties [8].
Ranjbar-Mohammadi et al. prepared a PLGA/Gum tragacanth (GT) blend via blending electrospinning and PLGA/GT core-shell nanofibers via coaxial electrospinning at various ratios [58]. The tetracycline hydrochloride (TCH), a hydrophilic model drug, was incorporated within these nanofibers in order to investigate the drug-release kinetics displayed by the produced membranes. The authors obtained uniform nanofibers of PLGA, blend PLGA/GT, and core-shell PLGA/GT [58]. Moreover, the drug release of the TCH from the blend nanofibers and the core-shell structures can be controlled throughout both membranes. However, the incorporation of the TCH into PLGA/GT core-shell nanofibers exhibited a more prolonged release for 75 days with a smaller burst release within the first 2 h [58]. Thus, the more sustained release of the TCH from the core-shell membranes allied with their proven biocompatibility, antibacterial, and excellent mechanical properties make them suitable for periodontal regeneration purposes [58].
In turn, Chachlioutaki et al. used silk sericin, a natural protein derived from silkworm cocoons, and PLGA to incorporate an anti-inflammatory drug, ketoprofen, by blending electrospinning [59]. Their results revealed an increase in the hydrophilic character of the composite membranes, good mechanical properties, and a sustaining drug release for up to 15 days [59]. Besides, the in vitro assays demonstrated that the silk sericin-PLGA composite membranes promote the attachment and proliferation of human gingival fibroblasts and induce a significant downregulation of the specific pro-inflammatory markers MMP-9 and MMP-3 and an upregulation of the anti-inflammatory gene IL-10 on lipopolysaccharide-simulated RAW 264.7 macrophages [59]. Further, Zupančič et al. incorporated resveratrol (RSV) at different quantities (1, 5, 10, and 20% (w/w)) into PCL electrospun nanofibers. RSV loaded into the PCL nanofibers below 5% resulted mostly in the form of solid dispersion, while at higher loading were observed nanocrystals on the fibers surface [60]. The PCL nanofibers loaded with RSV showed a bi-phase release kinetic due to the drug dissolution and cleavage of hydrogen bonding and hydrophobic interactions between the PCL and the RSV. Thus, the RSV-loaded PCL nanofibers showed their suitability to be used as drug-delivery systems for the treatment of periodontal disease [60].
In 2022, Zhong et al. produced via electrospinning a bi-layered nanofibrous membrane composed of an antibacterial layer of PLGA/Gelatin loaded with nano-silver (nAg) and an osteoconductive layer of PLGA/Gelatin loaded with nano-hydroxyapatite (nHA) for periodontal tissue regeneration and reestablishment [61]. Both nAg and nHA were successfully incorporated into the electrospun PLGA/Gelatin nanofibers. Additionally, in vitro assays indicated that the nanofibrous membranes displayed excellent cytocompatibility and the PLGA/Gelatin nHA-PLGA/Gelatin osteoconductive layer exhibited an improved osteogenic ability for human osteoblast-like cells (MG63), as confirmed by the good cell viability and the increased alkaline phosphatase (ALP) activity, respectively [61]. On the other hand, the nAg-PLGA/Gelatin antimicrobial layer of the bi-layered nanofibrous membranes presented an effective antimicrobial capability against S. aureus and E. coli [61]. Jenvoraphot et al. combined poly(l-lactide-co-ε-caprolactone) (PLLCL) with Tetracycline to produce electrospun membranes that promote periodontal regeneration and deliver an anti-inflammatory and antibiotic drug [62]. The release profile of TC from the electrospun PLLCL membranes was characterized by a controllable slow release of the drug from the nanofibers. Besides, the in vitro assays demonstrated that the electrospun PLLCL membranes containing TC promote the proliferation of human oral fibroblast (HOF) and human oral keratinocyte (HOK) cells and exhibit antibacterial properties [62].
Peng et al. produced coaxial magnesium oxide (MgO) NPs-incorporated PCL/Gelatin core-shell nanocellulose membranes for periodontal tissue regeneration [63]. The encapsulation of the MgO within the core-shell PCL/Gelatin nanocellulose nanofibrous membranes assured their sustained release and promoted the adhesion and proliferation of human periodontal ligament stem cells (hPDLSCs) due to the biocompatibility and hydrophilicity of the Gelatin shell layer [63]. Furthermore, in vitro osteogenic and antibacterial activities were enhanced in the PCL/Gelatin nanocellulose membranes containing MgO nanoparticles [63]. Additionally, Pouroutzidou et al. reported the incorporation of the Moxifloxacin-loaded Silica-based mesoporous nanocarriers (MSNs) in PLGA composite membranes produced through electrospinning [64]. The produced drug-loaded composite fibrous membranes presented a controlled and prolonged release profile, ensuring a desirable antibacterial activity against a wide range of periodontal pathogens, good hemolytic behavior, and a protective effect on the erythrocytes [64].
4.4 Vascular grafts tissue regeneration
The electrospun nanofibers have emerged as a promising alternative to produce vascular grafts due to their ability to control the mechanical properties, the fiber diameters, the porosity, and the pore size [16]. In addition, electrospinning is a highly versatile method and allows to produce of materials with different structures and forms, as well as combining the advantages of synthetic and natural materials, which are vital for tissue-engineered grafts. Moreover, the different electrospinning techniques offer control over the alignment of the nanofibers, which can result in the orientation of the cells in a specific direction, able to provide the anisotropy encountered in certain organs including blood vessels [16].
Yang et al. prepared PCL solutions containing rapamycin (RM) that were electrospun outside the decellularized vascular graft in order to produce an RM-loaded vascular graft [65]. A small-diameter RM-loaded vascular graft showed superior mechanical properties and a sustained drug-release profile, which is considered ideal for prolonged bioactivity. Moreover, in vivo assays demonstrated that, 12 weeks after implantation, the hybrid electrospun RM-loaded decellularized vascular grafts were able to significantly reduce intimal hyperplasia without impairing reendothelialization and M2 macrophage polarization [65]. Kuang et al. combined the electrospinning and freeze-drying technology to create core-shell structures of poly(l-lactide-co-caprolactone) (PLCL) (core) nanofibers coated with heparin/silk gel (shell) [66]. The core PLCL nanofibers were designed to provide mechanical support during vascular reconstruction, while the shell heparin/silk gel layer improved the biocompatibility of the grafts. The release of heparin from silk gel-PLCL composite nanofiber graft could regulate the microenvironment in the early stage after transplantation and inhibit intimal proliferation [66]. Besides, the graft showed an auspicious biodegradation rate and was safe. Furthermore, the composite nanofiber graft provided to be a useful model for remodeling vascular structures, and in vivo assays demonstrated that the graft remained unobstructed for a long period of time (for more than 8 months) [66].
In 2021, Han et al. produced coaxial electrospun core-shell structure fibers for small-diameter vascular grafts incorporated with puerarin (PUE) [67]. The gelatin by taking PLLA was used as a carrier material, while multi-walled carbon nanotubes (MWNTs) were used as reinforced materials. The obtained results revealed that the produced Gelatin/PLLA/MWNTs-PUE fiber membranes displayed wettability, degradability, and mechanical properties similar to those presented by the natural blood vessels [67]. Besides, the PUE release from the Gelatin/PLLA/MWNTs grafts occurred in an effectively prolonged and highly efficient manner. Moreover, the fiber membrane did not induce any cytotoxic effects or side effects on endothelial cells, as well as any hemolytic activity [67]. Likewise, Maleki et al. explored the electrospinning and freeze-drying methods to fabricate a bilayer vascular graft with a core-shell structure [68]. An anticoagulant core layer composed of electrospun nanofibers of silk fibroin (SF) and thermoplastic polyurethane (TPU) containing Heparin (Hep) and a shell highly porous hydrogel layer fabricated by the freeze-drying method was used to produce a bilayer tubular vascular graft [68]. Their results revealed that the electrospun TPU-SF-Hep fibers showed robust mechanical properties, while the hydrogel layer increased the viability of the smooth muscle cells (SMCs) [68]. The Hep displayed a sustained release from the proposed graft over 40 days, as well as excellent cell and blood compatibility, as shown by 3-(4,5-dimethylthiazol- 2-yl)-2,5-diphenyltetrazolium bromide (MTT) assay and platelet adhesion test [68].
4.5 Wound dressings
The electrospun nanofibrous membranes have been widely used for efficiently wound healing and skin regeneration due to their ability to mimic the randomly oriented 3D structure of the collagen fibers found within the skin’s ECM and suitable mechanical properties [8, 16]. Moreover, these membranes display high porosity, small and interconnected pore size, and high surface area, which provide a suitable microenvironment for cell adhesion and proliferation, as well as excellent nutrients/oxygen permeability, controllable evaporative water loss, and fluid drainage ability [8, 16]. Furthermore, various drugs, including antimicrobials, anti-inflammatories, growth factors, vitamins, and even cells, have been loaded into electrospun nanofibers for achieving antibacterial, antifungal, anti-inflammatory, antioxidant, analgesic, and anesthetic properties, which are vital for wound dressing applications [8, 16].
Yao et al. prepared peanut protein isolate (PPI)/PLLA nanofibers loaded with f 5, 10, and 15% of tetracycline hydrochloride (TCH) for being applied as wound dressing materials [69]. The PPI was developed using the alkali-dissolved acid precipitation method, and then the blend PPI/PLLA containing TCH was obtained via electrospinning. The incorporation of TCH did not affect the morphology of the nanofiber membranes, although the fiber diameter decreased when TCH content increased, while the wettability and mechanical properties improved [69]. Besides, the TCH was released from the PPI/PPLA composite nanofibers through a Fickian diffusion process. Moreover, the hemolysis rates were below 5%, which indicated the safety of the drug-loaded PPI/PLLA nanofibers, and the TCH-loaded PPI/PLLA composite nanofibers maintained the anticoagulant effect and showed low toxicity to cells [69]. The TCH/PPI/PLLA nanofibers also exhibited an inhibitory effect on S. aureus and E. coli growth and promoted skin wound healing in mice [69]. In turn, Yang et al. produced electrospun Janus nanofibers composed of PVP and EC loaded with a ciprofloxacin (CIP) and silver NPs (AgNPs) via a side-by-side electrospinning process as an effective antibacterial wound dressing [70]. The nanofibers presented a uniform appearance and cylindrical morphology with a clear Janus structure. The AgNPs were distributed on the EC side, while the drug was present in the PVP fibers [70]. In vitro assays demonstrated that over 90% of CIP was released within the first 30 minutes, confirming a strong antibacterial effect against both S. aureus and E. coli at the initial stages of the wound-healing process [70]. Lan et al. designed core-shell PVA/PCL nanofibers through coaxial electrospinning with dual release of tea polyphenols (TP) and ε-poly (L-lysine) (ε-PL) as antioxidant and antibacterial wound dressing materials [71]. The antioxidant TP incorporated in the PVA core showed a sustained release profile, while the antibacterial ε-PL in the PCL shell presented a fast release, allowing to achieve an antibacterial effect toward S. aureus and E. coli in the initial phase of the healing process and prolonged antioxidant activity [71]. Furthermore, the prepared coaxial core-shell nanofibers simultaneously incorporated with ε-PL and TP presented excellent cytocompatibility [71].
Zhong et al. fabricated through coaxial electrospinning PVA/PLA nanofibers embedded with Bletilla striata polysaccharide (BSP) and Rosmarinic acid (RA) to promote the wound-healing process [72]. Their results showed that the core-shell RA-BSP-PVA@PLA membranes were able to control the water vapor transmission rate (WVTR), and exhibited excellent flexibility, as well as better accommodate wounds [72]. Moreover, the MTT assay revealed that the RA-BSP-PVA@PLA nanofibers exhibited good biocompatibility and safety properties, as well as induced wound tissue growth in rat dorsal skin wound models and tissue sections [72]. Furthermore, the histological examination showed that the produced nanofibers enhanced the collagen deposition, the epidermal thickness, and the granulation tissue formation, as well as promoting the conversion of M1 macrophages into M2 macrophages, reducing the release of inflammatory factors, and promoting the occurrence of an effective wound-healing process [72].
Moreover, Mouro et al. incorporated Chelidonium majus L. into electrospun nanofibrous membranes composed of PCL, PVA, and Pectin (PEC) through emulsion electrospinning [73]. The electrospun PCL/PVA/PEC membranes exhibited suitable morphological, chemical, physical, and mechanical features for application as wound dressing materials [73]. Moreover, these membranes displayed excellent biological properties, namely antibacterial activity against S. aureus and P. aeruginosa, and the cytocompatibility assay provided no evidence of cytotoxicity in normal human dermal fibroblasts (NHDF cells) [73].
Recently, Sarviya et al. evaluated the performance of an ultrafine three-layer polymer nanofiber membrane produced via electrospinning [74]. In this approach, the first layer was produced with Polystyrene (PS) to act as a carrier layer and confer a suitable mechanical strength, the second layer consisted of PCL containing AgNPs endowing it with antibacterial properties, and finally, the third layer comprised of PEO act as a nonadhesive hydrophilic barrier layer with the potential to further support the healing process [74]. The cumulative Ag+ release was improved for a period up to 84 days and the findings highlight the importance to balance the antibacterial properties with the low toxicity in order to produce a biocompatible three-layer wound dressing material containing antibacterial properties [74].
4.6 Other textile-based applications
The electrospinning technique has allowed the creation of different textile materials with controllable diameters by changing several parameters, such as the solution properties, processing conditions, and environmental variables [8]. Besides, electrospun nonwoven membranes have displayed remarkable features, like high porosity, small pore size, high surface area, and good interconnected pore structure, which permits to efficiently capture ultrafine particles, like particulate and microbial contaminants [75, 76]. Moreover, these nonwoven structures are flexible, breathable, and comfortable, and therefore electrospinning has become a promising technique for producing PPE, particularly face masks [13, 75, 77].
In 2022, Geetha et al. dispersed Zinc Oxide (ZnO) NPs into a PVA/PVP polymer blend solution through electrospinning to produce efficient antimicrobial face masks [75]. The results obtained revealed that the produced ZnO/ PVA/PVP composite nanofiber displayed a pronounced antibacterial effect against different pathogenic bacterial strains [75]. In turn, Salam et al. incorporated different concentrations of Viroblock (VB) (0.5, 1.5, 2.5, 3.5, and 5.0%) into a Polyacrylonitrile (PAN)/ZnO solution via electrospinning [78]. The electrospun nanocomposite reached an antibacterial efficiency of 92.59% against S. aureus and 88.64% against P. aeruginosa, when the higher amount of VB (5.0%) was incorporated in the nanofibers, as well as a significant reduction in virus titer (37.0%) [78]. Hence, the VB-loaded PAN/ZnO nanofibers showed the potential to develop nanofiber-based personal protective equipment, such as facemasks and surgical gowns, due to their ability to kill enveloped viruses, such as coronaviruses and influenzas [78]. Further, Ferreira et al. produced an antibacterial face-mask filter composed of PCL combined with MgO and CuO NPs using an electrospinning technique [76]. The PCL filter dopped with CuO/MgO NPs exhibited structural stability up to 2 h of washing, offered filtering capacity, and additional antibacterial activity against two different bacteria strains, E. coli and S. aureus [76].
More recently, researchers started to explore the use of electrospun nanofibers as promise carriers for the local delivery of anticancer agents in order to circumvent some of the limitations presented by the simple nanostructures, like the low stability and the burst drug-release features, and effectively prevent cancer progression and improve their therapeutic index [79]. In fact, electrospun nanofibers have many advantages in cancer treatment due to their nanofibers size, surface modification, alignment variation, and drug molecules incorporation.
In this sense, Ahmady et al. investigated the release profile of capsaicin-loaded alginate (ALG) NPs from the PCL/chitosan (CS) nanofibers. Firstly, ALG NPs were prepared using different concentrations of cationic gemini surfactant and nanoemulsions as templates. Then, the optimized ALG NPs were loaded with 20 wt% capsaicin, a pharmacological natural agent with potent anticancer activity, into a blend of PCL and CS prepared at 2:1 volume ratio. The Cap-ALG NPs @ PCL/CS nanofibers were fabricated at a constant flow rate of 0.2 mL/h, using a working distance from the needle to the rotatory drum collector of 14 cm, with a rotation speed of 300 rpm, and a voltage of 16 kV [79]. The results obtained revealed a prolonged capsaicin release from 120 h to more than 500 h when capsaicin-loaded ALG NPs were incorporated into PCL/CS nanofibers. In addition, in vitro assays also demonstrated that The Cap-ALG NPs @ PCL/CS nanofibers could effectively inhibit the proliferation of MCF-7 human breast cells and did elicit any cytotoxicity effect on human dermal fibroblasts (HDF). Therefore, the long-term and controlled release of capsaicin from the electrospun nanofibers has a high potential for the prevention and treatment of cancer [79]. In turn, Seyhan et al. prepared PLA/ Polyethylene glycol (PEG) nanofibers loaded with amygdalin (AMG) and bitter almond kernel extract in order to prevent local breast cancer recurrence [80]. The produced electrospun PLA/PEG nanofibers containing AMG exhibited a sustained and controlled release extending up to 10 h. Additionally, the acquired in vitro data revealed that the nanofibers were able to induce cytotoxicity against MCF-7 breast cancer cells [80]. Therefore, AMG-loaded nanofibers arise as a highly promising approach for reducing the risk of local recurrence of cancer after surgery and can be directly embedded into solid tumor cells for treatment. Thus, electrospun nanofibers, particularly polymeric nanofibers, can be used to incorporate antitumor drugs for sustained and controlled release of chemotherapeutic compounds at a particular site over a specific period of time with low risks of toxicity and side effects to the healthy cells. Besides, the electrospun nanofibers have received increasing consideration as promising implantable textile-based materials for the on-site delivery of chemotherapeutic drugs in a sustainable release manner after surgical resection to inhibit tumor recurrence and prolong drug release at the tumor site [81].
Medical textiles are a field within the textile industry that has received increasing attention from researchers in recent years. Besides, the textile industry can benefit from nanotechnology, since it provides better functionalities and properties to the materials. Regarding that, tremendous efforts have been made to produce drug loading and delivery systems for tissue regeneration and wound healing able to release the drug in a controlled manner over the desired period of time.
Among the diverse methods used until now to manufacture nanoscale materials, electrospinning has been recognized as a viable technique to fabricate nanofibers from various types of textile materials that can act as carriers for the delivery of drugs due to their intrinsic properties, such as high surface area, high porosity, and small pore size. Moreover, the electrospun nanofibers have been loaded with a wide range of drugs using different electrospinning techniques, namely through blending, side-by-side, coaxial, tri-axial, emulsion, multi-needle, and needleless electrospinning. However, although the electrospinning techniques have the inherent benefits of simplicity, low cost, and versatility, all the techniques exhibit advantages and disadvantages depending on the polymeric solution and desired drug-release behavior. In fact, although the electrospun nanofibers produced by blending electrospinning are the most researched and applied for loading drugs by direct dissolving or dispersing in the polymeric solution; usually a rapid drug-release rate is observed. Thus, core–shell nanofibers produced by coaxial electrospinning and emulsion electrospinning have been highlighted to slow down the drug release. In addition, the coaxial and emulsion electrospinning provides an additional advantage for encapsulating fragile drug molecules, like growth factors, enzymes, and DNA, into the core protected by a shell, avoiding their denaturation. Further, these electrospinning techniques can involve the simultaneous use of two immiscible polymer solutions. Nonetheless, coaxial electrospinning needs a special syringe tip, while emulsion electrospinning requires the same basic setup as blending electrospinning. Moreover, side-by-side, tri-axial, and multi-needle electrospinning techniques offer the advantage of providing richer functionality from the different polymers and may result in biphasic and prolonged sustained release profiles. In turn, needleless electrospinning techniques addressed the limitations of traditional needle-based electrospinning and can produce high-throughput nanofibers using a wide variety of spinneret shapes and methods.
In addition, it is possible to tailor the properties of the electrospun textile materials by optimizing the parameters that influence the electrospinning (properties of the solution, processing variables, and environmental conditions), as well as the materials and configuration used in order to obtain the textile nanofibers with the desired morphology, size, orientation, porosity, wettability, mechanical properties, and degradation and drug-release kinetics for biomedical applications, including for bone, nerve, periodontal, and vascular tissue regeneration, wound dressings, and other textile-based materials, like PPE and cancer treatment.
However, until now, the data available concerning the nanofibers’ applications at the industrial scale are still very poor, being almost limited to the laboratory scale. Thus, preclinical and clinical assays are required for further validation of these materials and to help their transition to the market. Another future outlook with the potential to revolutionize various fields could be the application of electrospun nanofibers that exhibit changes in their physicochemical properties in order to respond to environmental stimuli, where the incorporation of drugs into smart electrospun nanofibers will enable precise activation-modulated or feedback-regulated control of the drug release. As research into this technology continues to progress, the potential applications for textile electrospun nanofibers will likely expand, ranging from drug loading and delivery systems to tissue regeneration and beyond.
Financial support was provided by the Portuguese Foundation for Science and Technology (FCT), I.P./MCTES through national funds (PIDDAC), in the scope of the FibEnTech Research Unit project (UIDB/00195/2020).
Invited author for book contribution—Isabel C. Gouveia; Pedagogical content and methodology—Isabel C. Gouveia and Cláudia Mouro; Scientific supervision and editing—Isabel C. Gouveia; Figures and tables layout—Cláudia Mouro.
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Written By
Isabel C. Gouveia and Cláudia Mouro
Submitted: 26 April 2023Reviewed: 07 August 2023Published: 21 December 2023
Open access peer-reviewed chapter
